Synthetic biomedical polymers can be broadly classified into non-biodegradable and biodegradable polymers. Those that are non-biodegradable are widely used when a medical device needs to be in place indefinitely or until such time as it is decided that the device is no longer required and can be safely removed, i.e. in permanent fixation devices. These polymers need to be completely non-biodegradable or have minimal degradation properties in the environment in which they are placed and are, for example, widely used in areas such as breast implants, in orthopedic applications such as bone fixation devices and, more recently, to replace important tissues such as heart valves. Polysiloxanes, polyurethanes and or their copolymers are widely used in such applications. Other examples include polyesters such as Dacron and polyetherether ketones (PEEK).
On the other hand, biodegradable polymers provide mechanical support and act as a platform for biological tissues to regenerate or repair when used in vivo and degrade after a period of time depending on the type of biodegradable polymer and the tissue environment. For these reasons, biodegradable polymers are particularly useful in orthopedic applications and tissue engineered products and therapies.
The vast majority of biodegradable polymers studied belong to the polyester family. Among these poly(α-hydroxy acids) such as poly(glycolic acid), poly(lactic acid) and a range of their copolymers have historically comprised the bulk of published material on biodegradable polyesters and have a long history of use as synthetic biodegradable materials in a number of clinical applications. Among these applications, poly(glycolic acid), poly(lactic acid) and their copolymers, poly(p-dioxanone), and copolymers of trimethylene carbonate and glycolide have been the most widely used. Their major applications include resorbable sutures, drug delivery systems and orthopedic fixation devices such as pins, rods and screws. Among the families of synthetic polymers, the polyesters have been attractive for use in these applications because of their ease of degradation by hydrolysis of their ester linkage, the fact that their degradation products are resorbed through metabolic pathways in some cases and their potential to be tailored in terms of their structure to alter degradation rates.
Synthetic polymers offer the advantage of tailoring the mechanical and other properties required for an intended application by choosing appropriate monomers and monomer combinations. Accordingly, various co-polymerisation methods have been developed to prepare polymers with a broad range of mechanical properties from the polyester family. Copolymers of lactic, glycolic acid, and ε-caprolactone are a few examples.
Most high modulus and high strength polymers are generally very brittle and have elongation at failure around 10% or less. Examples include poly(lactic acid), poly(glycolic acid), and polyanhydrides. Similarly, high modulus biostable polymers such as PEEK and ElastEon 4 (U.S. Pat. No. 6,437,073 B1) have failure elongations less than 50%.
Various strategies have been employed to improve the toughness of synthetic polymers. One of the main strategies employed is rubber toughening, incorporation of micron size rubber particle to brittle amorphous polymers [J Appl Polym Sci, 76, 1074 (2000)]. This approach has the disadvantage of having to incorporate a second polymer which is not desirable for many medical implant applications due to incompatibility and mechanical property mismatch. Furthermore, the improvement in elongation comes with a compromise in modulus and strength. Similarly, the incorporation of inorganic fillers increases modulus but with a compromise in elongation and strength, and such materials are not desirable for medical implant applications where high elongation is required.
High strength biodegradable polymers are sought after for applications such as vascular stents, fracture fixation implants and in other orthopedic applications such as spinal cages. High modulus but less brittle materials that can hold mechanical properties for a period of time until the repair process is completed are particularly sought after. For example, in coronary stents after the balloon expansion of the affected region of the blood vessel, the mechanical support of the stent is required for several months allowing sufficient time for the damaged vessel to repair. During this period the cellular growth around the stent takes place to rebuild the damaged vessel. The retention of material properties under physiological conditions (37° C., in vivo) is required for such applications. Accordingly, it is critical that the materials have high modulus, strength and elongation to prevent the implant from brittle failure for optimum performance of the implant in the biological environment.
Polyurethanes as a class of synthetic polymers offer advantages over other classes of polymers in designing materials with such properties. A wide range of polymers with a variety of properties ranging from elastomers to rigid materials can be prepared by selecting a suitable combination of reagents in various proportions. Diisocyanates, polyols and chain extenders are the three main reagents used to prepare polyurethanes. A high proportion of the diisocyanate and the chain extender generally yield rigid polyurethanes with high modulus and high strength. The polyurethane formed by reacting only the chain extender and the diisocyanate are generally very rigid with high modulus but very brittle and difficult to thermally process. For example, a polyurethane prepared from 4,4′-methylenediphenyldiisocyanate (MDI) and 1,4-butanediol (BDO) is very brittle due to its high crystallinity and melts above 210° C. [Polyurethanes Chemistry, Technology and Applications, Ellis Harwood p 118 (1993)]. Further, such materials have high modulus but due to brittleness have very limited applications.
The incorporations of fillers, rubber toughening, polymer blending have been reported in the polyurethane literature as methods to improve the toughness [see, for example, J Appl Polym Sci, 76, 1074, (2000); Polymer, 39, 865, (1998); Macromolecules, 30, 2876, (1997); J Appl Polym Sci, 63, 1335, (1997); J Appl Polym Sci, 63, 1865, (1997); WO 2006010278]. Results of these studies show that the increase in elongation to failure (increased toughness) always comes with a reduction in elastic modulus. In most cases the failure elongation was less than 5%.
In medical implants such as spinal cages and vascular stents the retention of the initial strength of the material is crucial for the proper function of the implant. Brittle materials will fail due to motion or other forces present in the biological system. Likewise, materials used for implants in bone fracture fixation not only should have sufficient mechanical strength to stabilise the fixation but also should retain the strength for a period of time ranging from weeks to months for proper healing of the damaged bone.
Accordingly, one object of this invention is to develop polymer compositions with high modulus, high strength and high elongation at failure for applications requiring load bearing capabilities. Desirably the compositions are biocompatible and capable of retaining initial mechanical properties under physiological conditions until the regenerated tissue structure can develop sufficient mechanical properties and subsequently the polymer degrades when no further mechanical support is required.